Noninvasive Body Fluid Stress Sensing

ABSTRACT

Electrochemical impedance-based label-free and rapid biosensor for select bodily fluid biomolecule levels. Monoclonal antibodies to of biomolecule such as Cortisol were covalently attached to a 16-mercaptohexadecanoic acid functionalized gold working electrode using zero-length crosslinkers N-(3-dimethylaminopropyl)-N-ethylcarbodiimide and 10 mM N-hydroxysulfosuccinimide. Cortisol was detected in phosphate buffered saline (simulated tear fluid) using a simple ferrocyanide reagent with a lower limit of detection of 18.73 pM and less than 10% relative standard deviation.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional PatentApplication No. 62/037,006, filed Aug. 13, 2014, and U.S. ProvisionalPatent Application No. 62/101,143, filed Jan. 8, 2015, which areincorporated herein by reference as if set forth in their entirety.

TECHNICAL FIELD

This disclosure relates to the field electrochemical sensing.

BACKGROUND

Over the past 40 years, chronic stress has been increasingly implicatedin a wide and growing variety of humanity's most lethal andlife-altering diseases. These include such severe conditions asdiabetes, Alzheimer's disease, heart attacks, depression, osteoporosis,and immunosuppression, as well as nonlethal but still unfortunateproblems like common colds, back pain, and even erectile dysfunction. Infact, scientific literature shows that stress affects life expectancy indeveloped countries more than genetics and behavioral factors such assmoking.

Given the enormous impact of stress on human life and health worldwide,there is great potential in measuring and treating stress on apopulation-wide scale. Although stress is often described as asubjective emotional state, medically it has important biochemical andphysiological effects. These effects that can be quantified, such asincreased levels of a group of certain hormones including theglucocorticoids and catecholamines. However, physiologicalconcentrations of these hormones, even when elevated, are oftenextremely low in tears, saliva and serum (38.9±15.5, 46.3±16.0, and489.7±177.4 nM respectively), making precise measurement a continuingtechnical challenge.

SUMMARY

A modified electrochemical sensor using a microfluidic tear fluidcapture system has been made to detect stress and/or trauma relatedbiomolecules, such as cortisol. Moreover, other bodily fluids such assaliva or blood may be utilized.

In one embodiment, monoclonal antibodies were covalently attached to a16-mercaptohexadecanoic acid functionalized gold working electrode usingzero-length crosslinkers N-(3-dimethylaminopropyi}-N-ethylcarbodiimideand 1 OmM N-hydroxysulfosuccinimide. Cortisol was detected in phosphatebuffered saline (simulated tear fluid) using a simple ferrocyanidereagent with a lower limit of detection of 18.73 pM and less than 10%relative standard deviation. The cortisol assay presented herein retainsa highly reproducible and ultralow level of detection in a label-freeand rapid response configuration with more than adequate sensitivity fortear cortisol measurement.

These and other aspects of the embodiments disclosed herein will beapparent upon reference to the following detailed description andfigures.

DESCRIPTION OF DRAWINGS

FIG. 1. A. Basic scheme of an embodiment of an apparatus withthree-electrode system including (a) Ag/AgCl Reference Electrode, (b)Sensing Well, (c) Sample, (d) Au Working Electrode, (e) GDE, and (f) PtCounter Electrode. Note that all materials are exemplary and can besubstituted for by other suitable materials. Additionally, amultiplexible electrochemical impedance spectroscopy (MEIS) system inoperable connection with the apparatus is schematically depicted. B.Sample with target (a) Cortisol is placed within the sensing well onsurface of covalently immobilized monoclonal antibody (MAb) on the goldworking electrode surface with MAb (b) immobilized to the gold surface(d) covalently with 16-MHDA (c) and EDC/NHS. Cortisol (a) target insample binds to the MAb.

FIG. 2. Nyquist plots of nine different MAb immobilized electrodes runin Cortisol target solutions at: (a) 0 pg/ml, (b) 1 pg/ml, (c) 5 pg/ml,(d) 10 pg/ml, (e)50 pg/ml, (f) 100 pg/ml, (g) 500 pg/ml, (h) 1000 pg/ml,(i) 5000 pg/ml, and (j) 10000 pg/ml in PBS buffer with 100 mM potassiumferrocyanide redox probe.

FIG. 3A depicts from the concentration gradient the calculations of (a)slope and (b) R-square (tightness of fit) that are made and plottedagain frequency to determine optimal frequency of detection.

FIG. 3B. Impedance at 1.184 Hz was used and plotted againstconcentration of Cortisol in PBS over physiological ranges and beyondshowing sensor dynamic range (n=3). A slope of 31.672 ohms/pg/mL isobserved with an R² of 0.9532 and 10% RSD at the highest concentrationvariance.

FIG. 4. A. A basic scheme of an embodiment of an apparatus withthree-electrode system including (a) Ag/AgCl Reference Electrode, (b)Sensing Well, (c) Sample, (d) Au Working Electrode, (e) GDE, and (f) PtCounter Electrode. Note that all materials are exemplary and besubstituted for by other suitable materials. Additionally, amultiplexible electrochemical impedance spectroscopy (MEIS) system inoperable connection with the apparatus is schematically depicted. B.Sample with target (a) Cortisol is placed within the sensing well onsurface of covalently immobilized monoclonal antibody (MAb) on the goldworking electrode surface with MAb (b) immobilized to the gold surface(d) covalently with 16-MHDA (c) and EDC/NHS. Cortisol (a) target insample binds to the MAb.

FIGS. 5A-5I. Detection of biomarkers in tear simulated fluid. Acalibration curve used in a cortisol device to correlate measuredimpedance to a concentration of cortisol, as well as plots showing thedetection of many different biomolecules by a device of the invention.

FIGS. 6A-6F show depictions of biomarker detection data in blood.

FIG. 7 depicts a summary of biomarker data.

FIG. 8. Cortisol interferents test results. The signal to noise ratioresult from an ELISA assay using IgG anti-cortisol antibody against theprovided standard, the usual cortisol gradient, and the testedinterferents at 200 pg/mL of each respective analyte is depicted.

FIG. 9 depicts a summary of stress biomarker data utilizing cyclicvoltammetry.

FIG. 10 depicts stress biomarker data using an amperometric technique.

FIG. 11 depicts stress biomarker data using the SWV (Square WaveVoltammetry) technique.

DETAILED DESCRIPTION

Though blood has historically been the standard diagnostic testingfluid, tear fluid has gained attention as a powerful sensing medium inrecent years for three major reasons. Firstly, the tear film contains avast number of biomarkers. Secondly, the relative ease in acquiring tearfluid compared to acquiring blood from the patient has made the tears anideal substitute for the blood in diagnostic testing. Finally, tearfluid, like saliva, is much less complex in composition than blood andcontains fewer proteins which might interfere with electrochemicalsensing.

Though there are some drawbacks to using tears (for example theavailable volume and target concentration are much less than those ofblood), these difficulties are outweighed by the benefits of easier andless invasive sampling and better sensor performance with lessbackground interference from non-target substances, proving the tearfilm to be the ideal diagnostic fluid for the stress sensor while stillcontaining measurable levels of cortisol.

Thus, in one aspect of the disclosure herein, a screen printedelectrode, an embodiment of which is shown in FIG. 1A-B, captures a tearsample via a novel microfluidic capture system that brings the sample tothe reagents and one or more molecular recognition units for cortisol(or other stress markers found in tears) encapsulated in the mesoporouscarbon inks of the sensor themselves has been developed using rapid,label-free and multiplexible electrochemical impedance spectroscopy(MEIS) that can be utilized at the point on care/injury. The molecularrecognition units may include one or more of antibodies, aptamers,peptides, synbodies, nucleic acids, tentacle probes, proteins, and thelike. Moreover, mesoporous carbon inks have been found to blockinterferents, leading to better test results.

Although stress is often described as a subjective emotional state, ithas been shown to have important biochemical and physiological effectswith dramatic impacts on human health. Consequently, monitoring stresslevels by sensing biochemical markers has the potential for making adramatic impact on stress management. Electrochemical impedancespectroscopy (EIS) is one such sensing method that has been successfulin label-free detection of a variety of extremely low concentrationtargets, including whole cell, protein biomarker, and small moleculetargets. Compared to other electrochemical methods, EIS has advantagesincluding speed (90 seconds per measurement), simplicity (no labelingrequirement as with “sandwich” assays) and sensitivity (detection ofpicomolar-concentration targets below the detection limits of many othermethods). This label-free sensing capability and ultralow detectionlimits make EIS an ideal sensing mechanism for cortisol in the tears.

EXAMPLE 1

A standard three-electrode system was used for impedance spectroscopymeasurements. The system is comprised of a Ag/AgCl reference electrode(CH Instruments, Austin, Tex.), a gold disk working electrode (GDE) (CHInstruments, Austin, Tex.), and a platinum counter electrode (CHInstruments, Austin, Tex.), with anti-cortisol antibodies(Sigma-Aldrich, St. Louis, Mo.) covalently attached to the workingelectrode surface to detect cortisol in the sample solution. A 1000 μLpipette tip (VWR International, Radnor, Pa.) was with the tip clippedwith a razor and fitted tightly over the GDE to create a plastic “well”able to hold around 0.2 mL of sample liquid. A diagram of this system isshown in FIG. 1.

Phosphate buffered saline (PBS) at pH 7.4 (EMD Biosciences, La Jolla,Calif.) was used to make all solutions unless otherwise noted. In orderto immobilize anti-cortisol antibody onto the surface of the gold diskelectrode (GDE), the GDE was first wet-polished with 120 figure-eightpasses on 3 μm aluminum oxide grit (CH Instruments, Austin, Tex.) andrinsed with distilled water. The 120 figure-eight polishing was thenrepeated with 1 μm and then 0.05 μm grit (CH Instruments, Austin, Tex.),after which the GDE was sonicated for 20 min in distilled water. Then,100 μL of a 1 mM 16-mercaptohexadecanoic acid (16-MHDA) (Sigma-Aldrich,St. Louis, Mo.) solution in reagent grade ethanol was placed into thesensing well and sealed in with Parafilm for 1 hr at room temperature.Next, the surface and sides of the GDE and sensing well were carefullyrinsed with distilled water. Control EIS measurements were performed onthe 16-MHDA-functionalized GDE using a “redox probe” of 100 mM potassiumferrocyanide (Sigma-Aldrich, St. Louis, Mo.) in PBS buffer to ensure anadequate and similar amount of MHDA was immobilized to each GDE. Thiswas determined by analyzing the impedance response of each individualGDE for comparability to one another.

Then, 100 μL of a PBS solution containing 40mMN-(3-dimethylaminopropyl)-N-ethylcarbodiimide (EDC) (PierceBiotechnology) and 10 mM N-hydroxysulfosuccinimide (sulfo-NHS) (VWRinternational) was placed in the sensing well. After 1 hr of incubationat room temperature, the electrode was rinsed with PBS buffer. Next, a100 μL droplet of a 10 μg/ml solution of anti-cortisol IgG (Aldrich) inPBS buffer was placed on the electrode and left at room temperature for1 hr, then rinsed off with PBS buffer. Finally, 100 μL of 1 mMethanolamine (Sigma-Aldrich, St. Louis, Mo.) in distilled water wasadded to the sensing well and incubated for 30 min at room temperatureto block all the unreacted carboxyl groups of the 16-MHDA and EDC/NHS.The electrode was then rinsed carefully with PBS buffer and stored inPBS at 4° C. until use.

Electrochemical impedance measurements were made using a CHI660CElectrochemical Workstation (CH Instruments, Houston, Tex.). Cortisol(Sigma-Aldrich, St. Louis, Mo.) sample concentrations from 0 to 10,000pg/mL (0 to 27.59 nM) were made in redox probe solution and stored at 4°C. until use. Each concentration of cortisol was then measured on eachof the antibody-immobilized electrodes.

For each measurement, 100 μL of the cortisol and redox probe solutionwas placed in the sensing well of the antibody-immobilized GDE. The ACpotential applied to the sample had an amplitude of 5 mV with a formalpotential (DC offset) of 150 mV, determined by a CV run on the bare(pre-immobilization) electrodes with redox probe. The AC voltage wasapplied at a range of frequencies from 1 to 100,000 Hz in 90 sec scanand the impedance magnitude and phase were recorded at each frequencyfor that sample. Real and imaginary impedances were calculated andplotted in a Nyquist plot for each sample. After each measurement, theGDE and sensing well were rinsed thoroughly with PBS prior to adding thenext sample.

For each electrode at each AC frequency tested, the impedance magnitudeat each cortisol concentration was correlated to log(concentration) witha slope and R² calculated. The impedance slopes and R² values were eachplotted against frequency in order to find the frequency which resultedin the best balance of high slope and R². The impedance values measuredat this “optimal” frequency were then used to generate the finalconcentration gradient allowing cortisol concentration to be estimatedfrom impedance.

AC sweeps of the bare, antibody immobilized, and biomarker (cortisol)bound electrodes yielded Nyquist plots. Nyquist plots of nine differentconcentrations of cortisol binding to one representative electrode areshown in FIG. 2. As the sample cortisol concentration increases, morecortisol binds to the antibodies on the electrode surface, thusincreasing the magnitude of the impedance at all frequencies and pushingthe Nyquist plots further out from the origin.

As expected, when the concentration of biomarker in vitro and thereforebound to the antibody increased, the impedance measured by the system(and therefore the signal) increased as well.

The slope and R² of a correlation plot of impedance versus targetconcentration change as a function of the AC potential frequency. Onerepresentative electrode is shown in FIG. 3a . Larger slope is desirableas it corresponds to a larger signal size (greater difference inimpedance values between different cortisol concentrations) which iseasier to measure with low proportional error. Larger R² is desirable asit indicates increased accuracy in the estimate of cortisolconcentration provided by the measurement. In FIG. 3a it can be seenthat R² is quite high for a range of frequencies below 100 Hz, but slopeis by far the best (largest) at very low frequencies and drops offrapidly as the frequency is increased. For all electrodes tested, theoptimal frequency to maximize slope was found to be 1.18 Hz. This is,therefore, the estimated optimal frequency at which thecortisol-antibody interaction is most effectively detected by EIS.

At this frequency (1.18 Hz), impedance data was compiled for severalsensors and plotted against corresponding concentrations to create theimpedance gradient shown in FIG. 3b . This gradient shows the impeccableaccuracy of this method in detecting the extremely low concentrations ofcortisol in the tear fluid. From the standard analytical definition oflower limits of detection (LLD), namely 3.3*slope divided by thestandard deviation, a LLD of 6.79 pg/mL (18.73 pM) was quantified inunder 90 sec detection time per sample. This clearly identifies measuredtear cortisol levels with a high degree of accuracy and a <10% sensor tosensor variance at any concentration.

The LLD of 18.73 pM is three full orders of magnitude below the typicalcortisol concentration range of around 40 nM in tears, which at firstglance seems to be a wildly excessive level of sensitivity for the tearsensing application. But in fact, this ultra-sensitive detection isprecisely what is needed to make this cortisol assay translatable fromthe laboratory to a physical real-world sensor device. A reproducibleand reliable sensor requires a low variance in not only theelectrochemical assay but also the physical device implementation. Tearsample sizes are unlikely to exceed 10 microliters in volume, as thereis just not that much tear liquid to collect, and some of the volumewill inevitably be lost by sticking to the walls of the sampling systemor other fluidics required to bring the sample from the eye surface tothe sensing electrodes.

As a result, ensuring a consistent reproducible volume of fluid in theelectrochemical sensing area (the functionalized electrodes) isextremely difficult without increasing the volume—diluting the targetconcentration by a known factor which can then be accounted for tocalculate the original sample concentration. Furthermore, tears do notcontain the high concentrations of redox mediators such as ferrocyanidewhich are needed for the electrochemistry to work. Thus, it would bedifficult for an actual sensor device to avoid diluting the 40 nM or soof cortisol with additional reagents, in order to both increase thetotal liquid volume to a workable amount that ensures a consistentvolume can reach the functionalized electrode area every time, andprovide a sufficient mediator concentration for the sensor'selectrochemistry. Because of this, a commercially viable tear cortisolsensor must provide reproducible measurements not in the 1-100 nM range,but rather in the range below 10 nM (for example, when a sample withabnormally low cortisol level is diluted even further by the deviceduring processing).

This is precisely what the EIS-based assay presented here allows. Withan LLD of below 0.02 nM, a 10 μL tear sample with 40 nM cortisol couldbe diluted 100× and still be well within the linear range of anEIS-based cortisol sensor. The cortisol assay shown here is thereforemore than capable of meeting the technical challenge of distinguishingamong low cortisol concentrations in tears.

In this work, the measurement of very low concentrations of cortisol isdemonstrated with reproducibility and high sensitivity using a simpleand label-free EIS-based biosensor. A replicated sensor set yieldedoptimal binding at 1.184 Hz with a reproducibility, at highestvariability under 10% relative standard deviation. The degree of fit wasmeasured to be 0.9532 with a responsivity of 31.672 ohms/pg/mL and alower limit of detection of 18.73 pM. This work shows that accurate andquick measurement of small changes in cortisol levels, even those as lowas typically found in human tear fluid, is technologically feasible,even after accounting for the practicalities of physical sensor designwhich may require further dilution of already low-concentration targetsfor reproducible performance.

In another aspect of the disclosure herein, a screen printed electrode,an embodiment of which is shown in FIG. 4, captures a body fluid samplevia a novel microfluidic capture system that brings the sample to thereagents and one or more molecular recognition units for cortisol (orother stress markers found in fluids) encapsulated in the mesoporouscarbon inks of the sensor themselves has been developed using rapid,label-free and multiplexible electrochemical impedance spectroscopy(MEIS) that can be utilized at the point on care/injury. While tearfluid is used in this embodiment, blood can also be used as shown inFIG. 6.

The molecular recognition units may include one or more of antibodies,aptamers, peptides, synbodies, nucleic acids, tentacle probes, proteins,and the like. Moreover, mesoporous carbon inks have been found to blockinterferents, leading to better test results.

EXAMPLE 2

While the following example is for detection of cortisol, similarprotocols are used for detection of other biomolecules of interest. Tearfluid or blood are used in this example but other bodily fluids may beused as well.

A standard three-electrode system was used for impedance spectroscopymeasurements. The system is comprised of a Ag/AgCl reference electrode(CH Instruments, Austin, Tex.), a gold disk working electrode (GDE) (CHInstruments, Austin, Tex.), and a platinum counter electrode (CHInstruments, Austin, Tex.), with anti-cortisol antibodies(Sigma-Aldrich, St. Louis, Mo.) covalently attached to the workingelectrode surface to detect cortisol in the sample solution. A 1000 μLpipette tip (VWR International, Radnor, Pa.) was with the tip clippedwith a razor and fitted tightly over the GDE to create a plastic “well”able to hold around 0.2 mL of sample liquid. A diagram of this system isshown in FIG. 4.

Phosphate buffered saline (PBS) at pH 7.4 (EMD Biosciences, La Jolla,Calif.) was used to make all solutions unless otherwise noted. In orderto immobilize anti-cortisol antibody onto the surface of the gold diskelectrode (GDE), the GDE was first wet-polished with 120 figure-eightpasses on 3 μm aluminum oxide grit (CH Instruments, Austin, Tex.) andrinsed with distilled water. The 120 figure-eight polishing was thenrepeated with 1 μm and then 0.05 μm grit (CH Instruments, Austin, Tex.),after which the GDE was sonicated for 20 min in distilled water. Then,100 μL of a 1 mM 16-mercaptohexadecanoic acid (16-MHDA) (Sigma-Aldrich,St. Louis, Mo.) solution in reagent grade ethanol was placed into thesensing well and sealed in with Parafilm for 1 hr at room temperature.Next, the surface and sides of the GDE and sensing well were carefullyrinsed with distilled water. Control EIS measurements were performed onthe 16-MHDA-functionalized GDE using a “redox probe” of 100 mM potassiumferrocyanide (Sigma-Aldrich, St. Louis, Mo.) in PBS buffer to ensure anadequate and similar amount of MHDA was immobilized to each GDE. Thiswas determined by analyzing the impedance response of each individualGDE for comparability to one another.

Then, 100 μL of a PBS solution containing 40mMN-(3-dimethylaminopropyl)-N-ethylcarbodiimide (EDC) (PierceBiotechnology) and 10 mM N-hydroxysulfosuccinimide (sulfo-NHS) (VWRinternational) was placed in the sensing well. After 1 hr of incubationat room temperature, the electrode was rinsed with PBS buffer. Next, a100 μL droplet of a 10 μg/ml solution of anti-cortisol IgG (Aldrich) inPBS buffer was placed on the electrode and left at room temperature for1 hr, then rinsed off with PBS buffer. Finally, 100 μL of 1 mMethanolamine (Sigma-Aldrich, St. Louis, Mo.) in distilled water wasadded to the sensing well and incubated for 30 min at room temperatureto block all the unreacted carboxyl groups of the 16-MHDA and EDC/NHS.The electrode was then rinsed carefully with PBS buffer and stored inPBS at 4° C. until use.

Electrochemical impedance measurements were made using a CHI660CElectrochemical Workstation (CH Instruments, Houston, Tex.). Cortisol(Sigma-Aldrich, St. Louis, Mo.) sample concentrations from 0 to 10,000pg/mL (0 to 27.59 nM) were made in redox probe solution and stored at 4°C. until use. Each concentration of cortisol was then measured on eachof the antibody-immobilized electrodes.

For each measurement, 100 μL of the cortisol and redox probe solutionwas placed in the sensing well of the antibody-immobilized GDE. The ACpotential applied to the sample had an amplitude of 5 mV with a formalpotential (DC offset) of 150 mV, determined by a CV run on the bare(pre-immobilization) electrodes with redox probe. The AC voltage wasapplied at a range of frequencies from 1 to 100,000 Hz in 90 sec scanand the impedance magnitude and phase were recorded at each frequencyfor that sample. Real and imaginary impedances were calculated andplotted in a Nyquist plot for each sample. After each measurement, theGDE and sensing well were rinsed thoroughly with PBS prior to adding thenext sample.

For each electrode at each AC frequency tested, the impedance magnitudeat each cortisol concentration was correlated to log(concentration) witha slope and R² calculated. The impedance slopes and R² values were eachplotted against frequency in order to find the frequency which resultedin the best balance of high slope and R². The impedance values measuredat this “optimal” frequency were then used to generate the finalconcentration gradient allowing cortisol concentration to be estimatedfrom impedance.

As expected, when the concentration of biomarker in vitro and thereforebound to the antibody increased, the impedance measured by the system(and therefore the signal) increased as well. See FIGS. 5A-5I throughFIG. 8.

In this work, the measurement of very low concentrations of cortisol isdemonstrated with reproducibility and high sensitivity using a simpleand label-free EIS-based biosensor. A replicated sensor set yieldedoptimal binding at 1.184 Hz with a reproducibility, at highestvariability under 10% relative standard deviation. The degree of fit wasmeasured to be 0.9532 with a responsivity of 31.672 ohms/pg/mL and alower limit of detection of 18.73 pM. This work shows that accurate andquick measurement of small changes in cortisol levels, even those as lowas typically found in human tear fluid, is technologically feasible,even after accounting for the practicalities of physical sensor designwhich may require further dilution of already low-concentration targetsfor reproducible performance.

Moreover, as summarized in FIG. 7, many biomolecules of interest can bedetected, such as cortisol, glucose, lactate, lactoferrin, IgE,catecholamines, 5-100 beta, neuron specific enolase, glial fibrillaryprotein, and tumor necrosis factor-alpha.

Turning to FIGS. 9-11, a summary of stress biomarker data is depicted.Cyclic Voltammetry (CV) is an electrochemical technique which measuresthe current that develops in an electrochemical cell under conditionswhere voltage is in excess of that predicted by the Nernst equation. CVis performed by cycling the potential of a working electrode, andmeasuring the resulting current. FIG. 9 shows a CV overlay of A EP, BNE, C DA, and D Cort (see structures in figure). Concentrations of DA,EP, Cort, and NE are 0.04M, 0.04M, 0.04M, and 0.1M respectively. Minimaloverlapping of signals is indicated by E, where F indicates largeoverlapping of signal peaks.

FIG. 10 depicts stress biomarker data from an amperometric technique.Amperometry in chemistry and biochemistry is the detection of ions in asolution based on electric current or changes in electric current.(Inlaid) An Amp-it of DA with the voltage applied at the oxidation peakof the CV, 0.52V, at A 2 sec, B 12 sec, C 20 sec during the AMP-it. Theouter graph is a calibration curve which plots current versusconcentration of DA at times (a), (b), and (c) during the AMP-it.Logarithmic fits of this calibration curve at different times A, B, andC have R² of 0.9566, 0.9547, and 0.9540 respectively.

FIG. 11 depicts the SWV (Square Wave Voltammetry) technique at 30 Hzused to determine concentration of EP v. current at the oxidation peak,0.23 V. The SWV technique at 20 Hz is used to determine concentration ofDA v. current at the oxidation peak, 0.22 V. The SWV technique at 20 Hzused to determine concentration of NE v. current at the oxidation peak,0.23 V. The SWV technique at 15 Hz is used to determine concentration ofCort v. current at the oxidation peak, 0.18 V.

The embodiments described above are not intended to be limiting.

What is claimed is:
 1. An electrochemical sensor, comprising: areference electrode and a counter electrode; a sensing well disposedbetween said reference electrode and said counter electrode; and afunctionalized working electrode disposed within said sensing well;wherein one or more of a molecular recognition unit to a stress markeris coupled to said functionalized working electrode.
 2. The sensor ofclaim 1, wherein said one or more molecular recognition units comprisemonoclonal antibodies are covalently attached to a16-mercaptohexadecanoic acid functionalized working electrode.
 3. Thesensor of claim 2, wherein said one or more monoclonal antibodies areattached to said functionalized working electrode with zero-lengthcrosslinkers N-(3-dimethylaminopropyi}-N-ethylcarbodiimide and 10 mMN-hydroxysulfosuccinimide.
 4. The sensor of claim 3, wherein saidreference electrode comprises a Ag/AgCl electrode.
 5. The sensor ofclaim 4, wherein said counter electrode comprises a Pt electrode.
 6. Thesensor of claim 1, further comprising a multiplexible electrochemicalimpedance spectroscopy system in operable arrangement therewith.
 7. Amethod for sensing a biomolecule in a bodily fluid, comprising:detecting, after a sensor having a molecular recognition unit to saidbiomolecule coupled to a working electrode is contacted with a bodilyfluid sample, an amount of said biomolecule bound to said molecularrecognition unit of said sensor using multiplexible electrochemicalimpedance spectroscopy.
 8. The method of claim 7, wherein said one ormore molecular recognition units comprise monoclonal antibodies arecovalently attached to a 16-mercaptohexadecanoic acid functionalizedworking electrode.
 9. The method of claim 7, wherein the fluid comprisestear fluid or blood.
 10. The method of claim 7, wherein said biomoleculeis one or more selected from the group of cortisol, glucose, lactate,lactoferrin, IgE, catecholamines, S-100 beta, neuron specific enolase,glial fibrillary protein, and tumor necrosis factor-alpha.